Transceiver phased arrays (also referred to as transceiver surface-coil arrays) may improve transmission performance (B1/√kW) and B1 homogeneity for magnetic-resonance head imaging up to 9.4 T (Tesla). To further improve reception performance and parallel imaging, the number of array elements has to be increased with correspondent decrease of their size. With a large number of small interacting antennas, decoupling is one of the most challenging aspects in the design and construction of transceiver arrays. Previously described decoupling techniques (for example, using geometric overlap, inductive or capacitive decoupling) have generally focused on eliminating only the reactance component of the mutual impedance. These decoupling methods may limit the obtainable decoupling by as much as −10 dB due to residual mutual resistance.
B1 magnetic field homogeneity generally refers to the homogeneity or inhomogeneity of a B1 magnetic field quantified by the standard deviation of the amplitude of the B1 magnetic field over a given region. The B1 magnetic field refers to the time-varying magnetic field generated by an RF antenna and is applied perpendicularly to the B0 magnetic field to alter the orientation of the nuclear spins of the nuclei of interest in the sample. The B0 magnetic field generally refers the primary static magnetic field applied by a MR system to a sample.
At high magnetic field strengths, where the object size becomes comparable to the RF wavelength, increased RF inhomogeneity, decreased transmit efficiency (μT/√W), and increasing local specific absorption rate (SAR) pose significant limitations for conventional single-channel transmit volume coils. For example, for body imaging, such limitations may be observed at 3 T and above, and for head imaging, such limitations may be observed at 7 T and above.
SAR generally refers to the measure of the rate at which energy is absorbed by the nuclei when excited to the B1 magnetic field. Transmission efficiency generally refers to the intensity of the B1 magnetic field generated by a coil element expressed in a unit of intensity (pT, μT, gauss, or Hz equivalent) as a function of the power applied to the RF coil to achieve that intensity.
To overcome these limitations, substantial effort has been focused on the development of transceiver phased arrays consisting of multiple independent (i.e. decoupled) RF antennas used simultaneously for both transmission and reception. Transceiver phased arrays provide improved homogeneity, enhanced transmit efficiency and decreased SAR through the use of RF shimming and parallel transmission.
RF shimming has been described in Adriany et al, Transmit and receive transmission line arrays for 7 Tesla parallel imaging. 53 MAGN. RESON. MED. 434-445 (2005); Mao et al, Exploring the limits of RF shimming for high-field MRI of the human head, 56(4) MAGN. RESON. MED. 918-922 (2006); Ibrahim et al, Insight into RF power requirements and B1 field homogeneity for human MRI via rigorous FDTD approach, 25(6) J. MAGN. RESON. IMAG. 1235-1247 (2007); Avdievich et al, Short Echo Spectroscopic Imaging of the Human Brain at 7T Using Transceiver Arrays, 62 MAGN. RES. MED. 17-25 (2009); and Kozlov et al, Analysis of RF transmit performance for a multi-row multi-channel MRI loop array at 300 and 400 MHz, Proceedings of the Asia-Pacific Microwave Conference, Melbourne, Australia 1190-1193 (2011).
Parallel transmission has been described in Katscher et al, Transmit SENSE, 49 MAGN. RESON. MED. 144-50 (2003); Zhu, Parallel excitation with an array of transmit coils, 51 MAGN. RES. MED. 775-784 (2004); and Zhang et al, Reduction of transmitter B1 inhomogeneity with transmit SENSE slice-select pulses, 57(5) MAGN. RESON. MED. 842-847 (2007).
Head arrays with surface coils as individual elements have been successfully utilized at 7 T and above. These efforts are described in Avdievich et al, Short Echo Spectroscopic Imaging of the Human Brain at 7T Using Transceiver Arrays, 62 MAGN. RES. MED. 17-25 (2009); Avdievich, Transceiver phased arrays for human brain studies at 7T, 41(2) APPL. MAGN. RESON. 483-506 (2011); Gilbert et al, A conformal transceive array for 7 T neuroimaging, 67 MAGN. RESON. MED. 1487-1496 (2012); and Shajan et al, A 16-Element dual-row transmit coil array for 3D RF shimming at 9.4 T, PROC. OF THE 20TH ANNUAL MEETING ISMRM, Melbourne, Australia 308 (2012).
Conventional MR systems typically employ a single transmission coil, to generate the RF magnetic field, commonly referred to as the B1 magnetic field. At higher magnetic field strengths, the wavelength of the RF magnetic field becomes comparable to the size of the sample (i.e. body imaging at 3T and above, head imaging at 7T and above). Also, with increase in the B0 magnetic field, the peak requirement for the B1 magnetic field has to also substantially increase while the transmission efficiency decreases. This result has been observed and described in “7 vs. 4T: F power, homogeneity, and signal-to-noise comparison in heat images,” Magn. Reson. Med. 46(1):24-30 (2001). Thus, at high magnetic field, it has been observed that a performance of a single transmission coil is significantly limited by increased RF inhomogeneity, decreased transmit efficiency (μT/W), and increased local specific absorption rate (SAR). Thus, transceiver phased arrays are more suited for higher magnetic field MR systems.
Transceiver phased arrays generally consist of multiple independent (i.e. decoupled) RF antennas configured to operate simultaneously for both transmission and reception. More array elements may improve the efficiency and homogeneity of the RF magnetic field being transmitted, reduce the effects of localized absorption regions, improve the sensitivity of the reception, as well as provide for parallel measurements. Each array element generally interacts with neighboring and non-neighboring elements in the array. The interaction is referred to as cross-talk and affects the RF field profile of the array, thus degrading the array transmission and reception performance, thereby lowering the signal-to-noise ratio (SNR) of the transceiver phased array. Examples of transceiver phased array are provided in US Patent Application, Publication No. 2012/0112748, titled “Transceiver Apparatus, System, and Methodology For Superior In-Vivo Imaging of Human Anatomy,” filed Aug. 18, 2011, by Hoby P. Hetherington, Jullie W. Pan, and Nikolai I. Advievich, which is incorporated by reference herein it is entirety.
Transceiver phased arrays may be used as conventional phased arrays for reception with the sensitivity of the receiver maintained. To provide sufficient coverage of the entire object during transmission and high signal-to-noise ratio (SNR) comparable with commercially available multi-channel receive-only arrays, the transceiver phased arrays may include multiple rows of smaller RF elements. For example, two or three rows of eight elements (2×8 and 3×8) may be employed for a head-sized array.
To overcome these limitations, substantial effort has been focused on the development of transceiver phased-arrays consisting of multiple independent (i.e. decoupled) RF antennas used simultaneously for both transmission and reception. With a large number of interacting RF antennas, decoupling, i.e. eliminating the cross talk, is becoming one of the most challenging and critical aspects in designing and constructing transceiver phased arrays.
It is known in the art to decouple array elements (also referred to as surface coils) of transceiver phased arrays using inductive or capacitive decoupling methodologies, thereby eliminating or reducing cross-talk. For certain geometries of individual antennas (e.g. overlapped surface coils) the cross-talk may include both reactive and a significant resistive components. All previously developed decoupling methods deal with eliminating only the reactive component of coupling (i.e. mutual inductance). Therefore, in these cases, use of any previously described decoupling schemes does not provide a complete decoupling of the array elements.
Compensating for both the resistive (real) and reactive (imaginary) components of the cross-talk may yield improved transceiver surface-coil arrays performance, thereby improving the sensitivity and imaging quality of the magnetic resonance system. Methods of decoupling the array elements using a resonant inductive decoupling circuit that eliminates the reactive component of the mutual impedance between array elements are also known in the art.
Overlapping of adjacent array elements is a common inductive decoupling technique and enables larger and greater numbers of RF coils to be used for a given circumference of the array. This technique is described in Roemer et al, The NMR phased array, 16 MAGN. RESON. MED. 192-225 (1990); Kraff et al, An eight-channel phased array RF coil for spine MR imaging at 7 T, 44(11) INVEST. RADIOL. 734-740 (2009); and Keil et al, Size-optimized 32-channel brain arrays for 3 T pediatric imaging, 66 MAGN. RESON. MED. 1777-1787 (2011), which are all incorporated by reference herein in their entirety.
Resonant inductive decoupling (RID) provides a way to compensate for both the reactive and the resistive components of the mutual impedance, Z12 (20). It also offers an easy way to adjust the decoupling, by changing the resonant frequency of the decoupling circuit through adjustment of a single variable capacitor. However, the placement and the geometry of these RID elements are critical since the RF field generated by the RID can significantly alter the RF field of the array.